KR102265599B1 - Enzyme immobilized glucose biosensor and the manufacturing method thereof - Google Patents

Abstract

The present invention includes a substrate including a plurality of electrodes and a porous matrix layer disposed on the substrate, wherein the porous matrix layer includes a glucose oxidase-coated nanofiber, bio for glucose measurement To a sensor and a method for manufacturing the same, and by using the biosensor of the present invention, it is possible to quickly and accurately measure the concentration of glucose in blood without affecting the hematocrit level.

Description

Enzyme immobilized glucose biosensor and manufacturing method thereof

The present invention relates to a biosensor for measuring glucose immobilized with glucose oxidase and a method for preparing the same.

Point-of-care (POC) blood glucose biosensors have been widely used as a major tool to manage the disease of diabetic neonates and critically ill patients. When used in stable adults, blood glucose meters are generally required to provide sensitive and accurate results. Although simple and relatively inexpensive, reporting inaccurate in a variety of clinical situations, such as sampling errors when using the same amount of blood used in a POC glucometer, the effect of hematocrit levels in patients with poor peripheral circulation, and changes in oxygen intensity [Weiner et al., 1991; Louie et al., 2000; Tang et al., 2000; Tang et al., 2001]. In particular, many studies have concluded that hematocrit fluctuations can significantly affect the accuracy of POC glucose measurements. This is because blood cells in the blood come into contact with the electrode surface and the electron-electrode reaction is recognized. To improve the glucose biosensor capability, an enzyme-based nanofiber layer was used in the biosensor to prevent adhesion of blood cells and increase assay sensitivity. Compared to other nanostructured materials, electrospun nanofibrous membranes have many advantages, such as high porosity, high surface area ratio, interconnectivity, and easy fabrication through convenient electrospinning technology. It has been widely used in the manufacture of biosensors requiring sensitive and rapid analysis.

Poly(vinyl alcohol) (PVA), a semi-crystalline polymer, has high hydrophilicity, chemical stability, low toxicity and good biocompatibility properties, such as hemodialysis membranes, artificial skin, and blood prosthetic devices. It has played an important role in the category of biomedical applied hydrogels.

Many strategies have been studied to reduce the effect of hematocrit on the glucose electrode. Forrow et al. tried to exclude erythrocytes using a porous carbon electrode impregnated with an activating reagent, but the analysis time was long, and there was a disadvantage that glucose in blood cells was oxidized, especially at high glucose concentrations.

The enzyme glucose oxidase (GOD, EC 1.1.3.4) converts glucose to gluconolactone, which is spontaneously hydrolyzed to gluconic acid with the simultaneous reduction of dioxygen to hydrogen peroxide. This oxidation reaction is a cofactor Accompanying the reduction of flavinadenine dinucleotide (FAD), FAD is responsible for the reduction of the enzyme and is tightly bound to the enzyme's polypeptide protein, but not covalently.

It is an object of the present invention, comprising a substrate including a plurality of electrodes and a porous matrix layer disposed on the substrate, wherein the porous matrix layer comprises a glucose oxidase-coated nanofiber, glucose measurement It is an object to provide a biosensor capable of quickly and accurately measuring a glucose concentration without affecting the hematocrit level in blood by providing a biosensor and a method for manufacturing the same.

However, the problems to be solved by the present invention are not limited to the problems mentioned above, and other problems not mentioned will be clearly understood by those of ordinary skill in the art from the following description.

A first aspect of the present invention provides a biosensor for measuring glucose, comprising a substrate including a plurality of electrodes and a porous matrix layer, wherein the porous matrix layer comprises nanofibers coated with glucose oxidase .

In an embodiment, the nanofiber is made of PVA/PAA, and the molar ratio of PVA and PAA may be 2:3 to 3:2.

In one embodiment, the glucose oxidase-coated nanofibers may have a diameter of 400 to 600 nm.

In an embodiment, the porous matrix layer may have a pore diameter of 2.5 to 5.0 μm before and after the sample is added.

In one embodiment, the thickness of the porous matrix layer may be 10 to 30㎛.

In one embodiment, in the biosensor, the glucose concentration measured in an environment of a hematocrit of 35% to 60% may be constantly measured within an error range of 5%.

In one embodiment, the glucose concentration constantly measured within the error range of 5% in the biosensor may be included in 37.1 mg/dL to 544.7 mg/dL.

In one embodiment, the biosensor may measure the glucose concentration within 1 to 3 seconds from the time the sample is added to the porous matrix.

A second aspect of the present invention comprises the steps of: preparing a substrate including a plurality of electrodes; preparing a hydrophilic polymer solution; Electrospinning the hydrophilic polymer solution on the substrate to prepare a porous matrix made of nanofibers; heat-treating the porous matrix (crosslinking process); and coating the porous matrix with glucose oxidase.

In an embodiment, the hydrophilic polymer may be PVA and PAA, and the molar ratio of PVA and PAA may be 2:3 to 3:2.

In one embodiment, in the electrospinning of the hydrophilic polymer solution on the substrate, the concentration of the hydrophilic polymer solution is 8 to 20 wt%, the voltage of the electrospinning is 10 to 80 kV, and the spinning speed is 5 to 30 ml/hr, and the spinning distance may be 10 to 30 cm.

In one embodiment, the coating of the porous matrix made of the nanofibers with glucose oxidase may be performed using a screen entrapment method, a bubble impregnation method, a spray method, or a dispensing method.

In an embodiment, the porous matrix layer may have a pore diameter of 2.5 to 5.0 μm before and after the sample is added.

The biosensor of the present invention, by including a porous matrix layer including enzyme-immobilized nanofibers, can quickly and accurately measure the blood glucose concentration without affecting the hematocrit level.

1 is a schematic diagram of a biosensor according to an embodiment of the present invention.
2 is a SEM photograph of nanofibers prepared by electrospinning with PVA having different degrees of hydrolysis.
3 is a graph showing the diameter of PVA according to the degree of hydrolysis.
4 is an SEM photograph of composite nanofibers according to the content of PVA/PAA.
5 shows cross-linking of nanofibers according to an embodiment of the present invention.
6 shows oxidation and reduction reactions of a biosensor according to an embodiment of the present invention.
7 shows a glucose oxidase coating method performed in a conventional biosensor.
8 shows a biosensor according to an embodiment of the present invention.
9 is a photograph showing the surfaces of a conventional biosensor and a biosensor according to an embodiment of the present invention, respectively.
10 is a graph showing a current graph with respect to blood glucose concentration, measured by a biosensor according to an embodiment of the present invention.
11 is a graph showing a current graph with respect to blood glucose concentration, measured by a biosensor according to an embodiment of the present invention.
12 is a graph showing a glucose concentration measured according to a hematocrit level, measured by a conventional biosensor.
13 is a graph showing a glucose concentration measured according to a hematocrit level, measured by a biosensor according to an embodiment of the present invention.
14 is a flowchart of a method for manufacturing a biosensor according to an embodiment of the present invention.
15 is a photograph of a step of cutting and attaching a matrix of a biosensor according to an embodiment of the present invention.

Hereinafter, embodiments will be described in detail with reference to the accompanying drawings. Like reference numerals in each figure indicate like elements.

Various modifications may be made to the embodiments described below. It should be understood that the embodiments described below are not intended to limit the embodiments, and include all modifications, equivalents or substitutes thereto.

Terms used in the examples are only used to describe specific examples, and are not intended to limit the examples. The singular expression includes the plural expression unless the context clearly dictates otherwise. In this specification, terms such as "comprise" or "have" are intended to designate that a feature, number, step, operation, component, part, or combination thereof described in the specification exists, but one or more other features It should be understood that this does not preclude the possibility of addition or existence of numbers, steps, operations, components, parts, or combinations thereof.

Unless otherwise defined, all terms used herein, including technical or scientific terms, have the same meaning as commonly understood by one of ordinary skill in the art to which the embodiment belongs. Terms such as those defined in commonly used dictionaries should be interpreted as having a meaning consistent with the meaning in the context of the related art, and should not be interpreted in an ideal or excessively formal meaning unless explicitly defined in the present application. does not

In addition, in the description with reference to the accompanying drawings, the same components are assigned the same reference numerals regardless of the reference numerals, and the overlapping description thereof will be omitted. In the description of the embodiment, if it is determined that a detailed description of a related known technology may unnecessarily obscure the gist of the embodiment, the detailed description thereof will be omitted.

Blood is classified into blood cells, which are cellular components, and plasma, which are liquid components. Blood cells are composed of red blood cells that carry oxygen, white blood cells that perform phagocytosis, and platelets that are involved in blood coagulation. Plasma contains 90-92% water, plasma protein It is composed of 7-8% and salts 1%, and other lipids, saccharides, and inorganic salts and non-protein nitrogen compounds, including urea, amino acids, and uric acid. The biosensor for measuring glucose measures the concentration of glucose contained in plasma in the blood. However, due to the interference of substances other than glucose in the blood, the conventional biosensor has a problem that it takes a long time to measure the concentration of glucose or the measured value may not be accurate.

1 is a schematic diagram of a biosensor according to an embodiment of the present invention.

1 , the biosensor of the present invention includes a substrate 100 including a plurality of electrodes and a porous matrix layer 200, and the porous matrix layer 200 is a glucose oxidase-coated nanofiber. includes

The electrodes are for measuring an electrochemical reaction, and are composed of a reference electrode and a working electrode for applying power, and, if necessary, may further include a pair of recognition electrodes for checking the amount of the sample required for analysis. have. That is, when a sample is identified in the pair of recognition electrodes, power is applied to the reference electrode and the working electrode to measure an electrochemical reaction. The number and structure of the electrodes may be changed as necessary.

The porous matrix layer is attached to the upper surfaces of the electrodes so as to cover the electrodes, so that blood can contact the electrodes after passing through the porous matrix layer, rather than directly contacting the electrodes. As blood passes through the porous matrix layer, it can filter hematocrit, which interferes with glucose measurement.

According to an embodiment of the present invention, the porous matrix layer includes nanofibers coated with glucose oxidase. Since the porous matrix made of nanofibers prepared through electrospinning can have a pore size that can be adjusted according to electrospinning conditions, glucose in the blood can be selectively permeated by using the porous matrix. In addition, the nanofibers constituting the porous matrix are coated with glucose oxidase, and the glucose oxidase is in contact with glucose in the blood to cause a redox reaction, and the biosensor electrode of the present invention has the resultant electron movement and current can be detected to measure the glucose concentration. In order to increase the contact area with blood, glucose oxidase is preferably coated on the surface of the nanofiber constituting the matrix. The coating layer on the surface of the nanofiber may include an electron transporter or a surfactant in addition to glucose oxidase, and the electron transporter may be ruthenium or peroxyanide.

In one embodiment of the present invention, the nanofiber is made of PVA/PAA, and the molar ratio of PVA and PAA may be 2:3 to 3:2. PVA/PAA is made of hydrophilic polymers, and unlike polymers mainly used in the prior art, crosslinking is possible even by simply using heat without the use of a crosslinking agent, so that hydrophilic hydrogel properties can be imparted. In particular, PVA/PAA prepared at an appropriate molar ratio has superiority in that it maintains porosity even at the maximum amount of water absorption and can maintain the cut-off function and nanoscale characteristics even in a state of immersing moisture. made up of nanofibers.

On the other hand, since PVA is derived from the hydrolysis of poly(vinyl acetate: PVAc), the properties of PVA are affected by the degrees of hydrolysis (DH). In particular, hydrolysis between 85 and 90% PVA with a degree of decomposition has lower mechanical properties and water resistance than PVA with a degree of hydrolysis between 98 and 99.9%, but is easily soluble in water and has excellent electroradiation (see Example 2).Therefore, the biosensor matrix of the present invention In order to form pores smaller and denser by reducing the diameter of the nanofibers constituting the layer, PVA having a degree of hydrolysis of 85 to 90 is preferably used, and most preferably PVA having a degree of hydrolysis of 88 may be used.

However, since PVA/PAA is bonded through ester formation between the -OH group of PVA and the -COOH group of PAA, there is a problem that it may be hydrolyzed in contact with moisture. In order to solve this problem, the PVA/PAA of the present invention was electrospun at a PVA:PAA molar ratio of 2:3 to 3:2, and cross-linked through a simple heat treatment process after electrospinning to increase water resistance (implementation) see Example 2 and FIG. 4). When the molar ratio of PAA to PVA is less than 0.25 or more than 0.75, the morphology of the fiber is broken, and it is difficult to maintain porosity and there may be problems that it does not perform the function as a nanofiber properly (Example 2 and FIG. 4 Note)

In one embodiment of the present invention, the glucose oxidase-coated nanofibers may have a diameter of 400 to 600 nm. The PVA/PAA fiber diameter before glucose oxidase coating was on average about 300 to 400 nm, but glucose oxidase was coated relatively densely and uniformly on the nanofiber surface in the form of crystals, increasing the overall diameter by about 100 to 200 nm make it When the diameter of the glucose oxidase-coated nanofibers is less than 400 nm, the surface area is increased, but there is a problem in that productivity is lowered.

In an embodiment of the present invention, the pore diameter of the porous matrix layer may be 2.5 to 5.0 μm. In order to exclude hematocrit, which interferes with the measurement of blood glucose concentration, the pores of the porous matrix are preferably 5.0 μm or less. In addition, when the size of the pores is less than 2.5 μm, blood permeation does not occur smoothly, and the measurement time may be delayed. More preferably, the pore diameter of the porous matrix layer may be 2.5 to 3.0 μm, and in order to more completely exclude hematocrit, which interferes with the measurement of glucose concentration in blood, the pores of the porous matrix are 3.0 μm or less. may be more preferable.

In addition, in the porous matrix layer, the pore diameter before and after the sample is added may be maintained in a range of 2.5 to 5.0 μm. That is, even if a sample such as blood is injected into the porous matrix layer, the nanofibers constituting the porous matrix layer of the present invention have high water resistance and are not hydrolyzed, so that the morphology of the fibers can be almost completely maintained (Example 2 and FIGS. see 4). Therefore, the biosensor of the present invention can measure the glucose concentration uniformly and accurately due to the porous matrix maintaining a uniform pore size even after a sample is added.

In one embodiment of the present invention, the thickness of the porous matrix layer may be 10 to 30㎛. When the porous matrix layer is more than 30 μm, the thickness of the porous matrix layer is preferably 30 μm or less, because glucose may be trapped in the porous matrix layer or filtered with hematocrit. The effect of excluding hematocrit may be reduced by increasing the probability of escaping hematocrit.

In one embodiment of the present invention, the biosensor may measure the glucose concentration measured in an environment of a hematocrit of 35% to 60% within an error range of 5%. After varying the hematocrit concentration in samples having the same glucose concentration, as a result of measuring glucose with the biosensor of the present invention, the measured glucose value was constant even when the hematocrit concentration was changed from 35% to 60%. (See Example 9 and FIG. 13)

In one embodiment of the present invention, the glucose concentration constantly measured within the error range of 5% in the biosensor may be 37.1 mg/dL to 544.7 mg/dL.

In an embodiment of the present invention, the biosensor may measure the glucose concentration within 1 to 3 seconds from the time the sample is introduced into the porous matrix. Conventional glucose biosensors require an average measurement time of 5 seconds or longer because the effect of hematocrit on the measured glucose concentration is corrected. However, since the biosensor of the present invention measures the glucose concentration after excluding hematocrit with the porous matrix without a calibration process, it takes only 1 to 3 seconds to measure the concentration.

14 is a flowchart of a method for manufacturing a biosensor according to an embodiment of the present invention.

Referring to FIG. 14 , the method for manufacturing a biosensor for measuring glucose of the present invention includes preparing a substrate including a plurality of electrodes ( 110 ); Preparing a hydrophilic polymer solution (120); Electrospinning the hydrophilic polymer solution on the substrate to prepare a porous matrix made of nanofibers (130); heat-treating the porous matrix (140); and coating the porous matrix with glucose oxidase (150).

In one embodiment of the present invention, the hydrophilic polymer is PVA and PAA, and the molar ratio of PVA and PAA may be 2:3 to 3:2.

Before the step of coating the porous matrix made of the nanofibers with glucose oxidase, the step of preparing an enzyme mixture containing glucose oxidase may be further included. The enzyme mixture may include an electron transporter or a surfactant in addition to glucose oxidase. The electron transporter may be ruthenium or ferrocyanide.

The heat treatment of the porous matrix 140 may be performed at a temperature of 60 to 200° C. for 10 to 120 minutes. When the heat treatment process is performed at less than 60 ℃, crosslinking may not be properly formed, and when it is carried out at a temperature of 200 ℃ or more, the polymers in the nanofiber may be melted or decomposed. In addition, when the heat treatment process is carried out for less than 10 minutes, a sufficient reaction does not occur and partial crosslinking may not occur. When the heat treatment process is performed for more than 120 minutes, the heat treatment process needs to be performed because the reaction has already sufficiently occurred. there is no

In one embodiment of the present invention, in the electrospinning of the hydrophilic polymer solution on the substrate, the concentration of the hydrophilic polymer solution is 8 to 20 wt%, the voltage of the electrospinning is 1 to 80 kV, The spinning speed may be 5 to 30 ml/hr, and the spinning distance may be 10 to 30 cm.

In one embodiment of the present invention, the step of coating the porous matrix made of the nanofibers with glucose oxidase may be performed using a screen entrapment method, a bubble impregnation method, a spray method, or a dispensing method, Any method known in the art as a matrix coating method may be used without limitation.

delete

delete

The biosensor manufacturing method of the present invention may further include cutting the porous matrix and attaching it to the electrode.

15 is a photograph of a step of cutting and attaching a matrix of a biosensor according to an embodiment of the present invention. Referring to FIG. 15 , after the porous matrix film prepared using electrospinning and coated with glucose oxidase is cut along the red line, it can be attached to the substrate on which the electrode pattern is formed.

<Example 1: Preparation of Materials>

All structural materials for the sensor chip and reactants are commercially available. Polyvinyl alcohol (PVA) having a degree of polymerization (DP) of 1,700 and a degree of hydrolysis (DH) of 88%, 96%, 99.9% is manufactured by Kuraray Co., Ltd. Ltd., Japan. Poly(acrylic acid; PAA) having a molecular weight of 250 kDa was obtained from Wako Pure Chemical Industries, Ltd., Japan Glucose oxidase (GOD, EC 1.1.3.4,15,000-25,000 units g- ¹ Type II from Aspergillusniger ), analytical reagents with a purity of 99% ruthenium (III) chloride, Triton X-100 wetting agent and β-d-(+)-glucose were purchased from Sigma-Aldrich. Used to prepare samples, buffers, and enzyme solutions, 0.1M phosphate buffer solution (PBS) was _{prepared as a mixed solution of KH 2} PO ₄ and K ₂ HPO _4. The pH of the prepared solution was 7.03. The enzyme solution was mentioned above It was prepared by dissolving in one buffer solution.

<Example 2: PVA hydrolysis degree and PVA/PAA content ratio>

After preparing PVA having a degree of hydrolysis of 88%, 92%, 96%, and 99.9%, respectively, they were electrospun under the same conditions to measure the diameter.

2 is a SEM photograph of nanofibers prepared by electrospinning with PVA having different degrees of hydrolysis. Figure 2a is a PVA nanofiber having a degree of hydrolysis of 88%, Figure 2b is a PVA nanofiber having a degree of hydrolysis of 92%, Figure 2c is a PVA nanofiber having a degree of hydrolysis of 96%, Figure 2d is a PVA nanofiber having a degree of hydrolysis of 99.9% This is an SEM picture of the fiber. It can be seen that a hydrolysis degree of 88% forms the thinnest fiber.

3 is a graph showing the diameter of PVA according to the degree of hydrolysis. The average diameter of PVA nanofibers with a degree of hydrolysis of 88% is 190 nm, the average diameter of PVA nanofibers with a degree of hydrolysis of 92% is 200 nm, the average diameter of PVA nanofibers with a degree of hydrolysis of 96% is 220 nm, and the degree of hydrolysis is 99.9% The average diameter of phosphorus PVA nanofibers was 470 nm.

The degree of hydrolysis of PVA/PAA composite nanofibers according to the content ratio of PVA/PAA was confirmed. When the molar ratio of PAA to PVA (φPAA) was 0.25, 0.5, and 0.75, composite nanofibers were prepared by electrospinning, and then, the water resistance and hydrolysis degree of each nanofiber were observed.

4 is an SEM photograph of composite nanofibers according to the content of PVA/PAA. (a) of the first row (upper row) of FIG. 4 is a nanofiber in the case of φPAA=0.25, (b) is a nanofiber in the case of φPAA=0.5, (c) is a nanofiber in the case of φPAA=0.75 to be. The picture in the second row (bottom row) of FIG. 4 is an SEM picture taken after each nanofiber was immersed in water for 1 hour. (a) is a nanofiber in the case of φPAA=0.25, (b) is a nanofiber in the case of φPAA=0.5, and (c) is a nanofiber in the case of φPAA=0.75. In the case of φPAA=0.5, the fiber morphology was most completely preserved, and it was confirmed that the water resistance was good.

<Example 3: Preparation of PVA/PAA cross-linked nanofibers>

To prepare PVA/PAA cross-linked nanofibers, PVA is first dissolved in water to prepare a solution for electrospinning with a concentration of approximately 10% by weight. To improve the solubility of the PVA polymer, the PVA solution is heated to 80° C. in bulk to break the strong internal bonds that may be present in the PVA polymer. A PVA/PAA solution for electrospinning was prepared by mixing a PAA solution with a PVA solution, the molar ratio was φPAA=0.50, and the concentration of the PVA/PAA solution was adjusted to approximately 10 wt%. A PVA/PAA mixed solution is electrospun on a rotating metal collector under the same conditions. The PVA/PAA mixed nanofibers are heat-treated at 160° C. for 30 minutes, and an ester is formed between the -OH group of PVA and the -COOH group of PAA and chemically cross-linked.

5 shows cross-linking of nanofibers according to an embodiment of the present invention.

<Example 4: Preparation of glucose sensor>

The enzyme solution was prepared by mixing 1.4 g of GOD and 600 mg of ruthenium (III) chloride in 20 mL of PBS. Enzymes are immobilized on PVA/PAA using a physical absorption method. The prepared sensor was dried at room temperature for 24 hours, and stored for about 1 week for enzyme stabilization and maturation before measurement.

Electrochemical measurements were evaluated with a sensor chip consisting of a two-electrode system and performed using an electrochemical analyzer. The devices introduced included a cyclic voltammeter, an amperometer (WonATech Co., Ltd), and WPCIPG software version 1.0. A Ni/Cu electrode was installed on a polyethylene terephthalate (PET) substrate manufactured by Jaein Circuit Co., Ltd, Korea. All electrochemical devices were performed in a cell containing 0.25 L of 0.1 M phosphite buffer solution (PBS, pH 7.03) using Cu/Ni, Cu/Ni/Pt, AgCl, and carbon electrodes as reference electrodes at room temperature.

<Example 5: Cu/Ni electrode>

Cu/Ni electrodes were fabricated using a heat treatment to laminate a 20 μm thick copper film on a 100 μm thick PET substrate. Then, an ultraviolet (UV) photosensitive dry film was attached by a laminating method. A reverse photo mask was used to form the copper electrode pattern by UV exposure and wet etching. The copper electrode is metallic and not treated with water, but the oxygen in the air will react slowly at room temperature to form a copper oxide layer on the copper metal. The copper electrode is followed sequentially by electroless plating of nickel.

<Example 6: Measurement method>

SEM images were measured with a Hitachi S4800 scanning electron microscope at an accelerating voltage of 5 kV. Samples were coated with Pt. Fiber diameters were measured using Image J software. For each sample, the fiber diameter was measured at a factor of 50 and averaged. Electrochemical experiments were performed with a three-electrode system on an IM6ex electrochemical workstation (Zahner, Germany). Cyclic voltammetric measurements were performed in an unagitated electrochemical cell.

6 shows oxidation and reduction reactions of a biosensor according to an embodiment of the present invention.

Quantitative measurement of glucose in whole blood using an electrochemical biosensor starts when a drop of blood is introduced into the top and side surfaces of the test strip. The electrons produced by the reaction create an electric current. An electric current is generated by electrons produced during glucose oxidation. The current is metered to measure the glucose concentration in the whole-blood sample. Whole-blood glucose concentration was checked using a YSI 2300 (Yellow Springs Instrument Inc., Yellow Springs, OH) standard glucometer. A YSI standard glucose analyzer is used to measure glucose concentrations in duplicate with two glucose channels using glucose oxidase. Linearity ranges from 0 to 1000 mg/dL. The YSI analyzer self-weighs at intervals of every 15 minutes or after 5 measurements. The hematocrit level of each blood sample was measured by centrifuging the sample at 10,000 rpm for 5 minutes with a micro-capillary centrifuge (Model MB, International Equipment Company, Needham Heights, Mass.).

<Example 7: Morphology of PVA/PAA cross-linked nanofibers containing GOD>

To prevent dissolution, PVA/PAA composite nanofibers were successfully prepared using the heat treatment method. The formation of a chemically crosslinked network through the formation of an ester bond between the hydroxyl group of PVA and the carboxyl group of PAA ^{was clearly demonstrated through the analysis of 13} C-NMR spectrum and FT-IR in the solid state.

By preparing two types of samples (the sample in which the GOD solution was directly dispensed on the Cu/Ni electrode, and the sample in which the electrospun PVA/PAA nanofiber membrane was coated with the GOD solution), the hematocrit effect of the PVA/PAA nanofiber membrane was evaluated.

7 shows a glucose oxidase coating method performed in a conventional biosensor, and FIG. 8 shows a biosensor according to an embodiment of the present invention.

9 is a photograph showing the surfaces of a conventional biosensor and a biosensor according to an embodiment of the present invention, respectively.

9 (a) and (b) are SEM images of a case in which a glucose oxidase complex is plasma-treated on a PET film according to a conventional method, wherein GOD aggregates into large, non-uniform clusters on the PET film. could confirm that it had been On the other hand, (c) and (d) of FIG. 9 show SEM images after coating the PVA/PAA cross-linked composite nanofibers by dipping them in a glucose oxidase solution, and then drying in a vacuum.

Referring to FIG. 9(c), according to an embodiment of the present invention, when the PVA/PAA nanofiber film is coated by dipping it in a GOD equivalent solution, the GOD is relatively uniformly deposited on the PVA/PAA nanofiber. could confirm that In addition, the fiber diameter was increased from approximately 350 ± 40 nm to 510 ± 50 nm after treatment with the GOD solution, and it was observed that GOD was uniformly formed on the PVA/PAA nanofibers. In the higher-resolution SEM photograph (FIG. 9 (d)), it was confirmed that the GOD particles formed on the PVA/PAA nanofibers were relatively evenly deposited on the nanofiber surface as nanocrystals within the structure. GOD particles were deposited on PVA/PAA nanofibers, increasing the overall diameter by approximately 100-150 ± 20 nm, and were successfully immobilized on PVA/PAA nanofibers without agglomeration.

<Example 8: Electrochemical ability of PVA/PAA-GOD coated nanofibers>

The electrochemical reaction of the PVA/PAA-GOD-coated nanostructures with respect to Cu/Ni was confirmed using cyclic voltammetry.

10 is a graph showing a current graph with respect to blood glucose concentration, measured by a biosensor according to an embodiment of the present invention. 10 is a low electrical resistivity (0.01 Ω or less) of the cyclic voltam response of the PVA/PAA-GOD-coated nanostructure layer on the Cu/Ni/PET film, which is a biosensor electrode having a uniform electrical resistance. suggests that it is possible to produce The low electrical resistivity of the sensor is advantageous because reduction peaks occur more rapidly than with high resistivity electrodes.

The reproducibility of the PVA/PAA-GOD coated nanofiber glucose biosensor was evaluated. The biosensor was tested using the amperometric method.

11 is a graph showing a current graph with respect to a blood glucose concentration measured by a biosensor according to an embodiment of the present invention, and shows a change in the reaction current at each whole blood glucose concentration. Referring to FIG. 11 , the PVA/PAA-GOD-coated nanofibers on the Cu/Ni electrode had a linear correlation coefficient ( R ² ) of 0.9924 and showed a constant current of 3 or less. The values observed in FIG. 11 were measured in 10 repetitions of disposable PVA/PAA-GOD-coated nanofibers on a Cu/Ni electrode glucose biosensor. As a result, the disposable Cu/Ni electrode glucose biosensor exhibited linearity and significant reproducibility. was shown. For the PVA/PAA-GOD coated nanofiber Cu/Ni electrode, the overall nonlinear error was 2.15% (coefficient of variation (CV, %) = standard deviation/arithmetic average Ⅷ 100). According to ISO 15197 [ISO 15197 First edition, 2003], an acceptable production criterion for a blood glucose sensor for clinical application is a non-linear error rate within 7%.

Whole blood samples with 325 and 450 mg/dL glucose concentrations were measured using the biosensor of the present invention at a scan rate of ^{0.1 Vs −1 .} The voltam profile shows characteristic hydrogen absorption/desorption curves according to Cu/Ni oxidation and reduction peaks (labeled), confirming the presence of PVA/PAA-GOD-coated nanofibers on Cu/Ni electrodes. Glucose was oxidized by electrochemical transformation of ruthenium chloride. Ruthenium chloride was continuously regenerated by the electrode under oxidizing conditions. The magnitude of the reaction current depends on the glucose concentration. Reaction currents were obtained from CV results (approximately +0.065 V) run at whole-blood glucose concentrations after blood samples were injected into PVA/PAA-GOD coated nanofibers. PVA/PAA-GOD coated nanofibers on Cu/Ni/PET films formed biosensor electrodes with very low electrical resistivity. Since the carbon electrode has a higher electrical resistivity (approximately 80 Ω to 1 kΩ ± 5 to 10%) than the Cu/Ni electrode (approximately 0.01 Ω) and requires a high applied voltage (approximately 0.2 to 0.7 V), the Enzymatic reactions require longer reaction times to reach a steady state. In comparison, the PVA/PAA-GOD coated nanofiber Cu/Ni electrode reaches a stable state within 3 seconds.

<Example 9: Hematocrit effect of PVA/PAA-GOD coated nanofibers>

For highly sensitive glucose biosensors, separation of plasma-containing glucose based on the size-exclusion principle is efficient. In human blood, red blood cells have a diameter of 6 to 9 μm and a thickness of 1.8 to 2.8 μm, and white blood cells have a diameter of 6 to 10 μm or more. Thus, a structure with a cut-off size of about 5 μm can tolerate glucose in the plasma, instead of blocking most of the white blood cells and red blood cells. Since the nanofiber of the present invention can control the size of the pores by controlling the conditions of electrospinning, it is not affected by blood cells and has high sensitivity to glucose.

Regarding the hematocrit effect, the biosensor of the present invention measured changes in the glucose concentration (140, 230, and 309 mg/dL) according to the hematocrit concentration (35, 42, 50, 60%).

12 is a graph showing the glucose concentration measured according to the hematocrit level, measured by the conventional biosensor, and FIG. 13 is the glucose measured according to the hematocrit level, measured by the biosensor according to an embodiment of the present invention. Concentration graph is shown. 12 and 13 are graphs of measurement results of glucose concentrations ranging from 37.1 mg/dL (2.06 mmol/L) to 544.7 mg/dL (30.24 mmol/L) at a hematocrit level of 35% to 60%.

Referring to FIG. 12 , as the hematocrit ratio increased, the current of the glucose biosensor decreased because red blood cells and white blood cells were attached to the electrode surface. According to the Cottrell equation,

I = (nFAD ^1/2 C) / (πt) ^1/2

(where I = current, n = number of electrons moving in the reaction (for ruthenium chloride, n = 1), F = Faraday constant (charge of 1 mole of electron = 96,485 Coulombs/mol), A = electrode area (cm) ² ), D = diffusion coefficient, C = reaction concentration (mol/cm ³ ), t = time (seconds)), efficient reduction of the electrode surface leads to low current values.

However, referring to FIG. 13 , since the biosensor current of the present invention tested under the same conditions was not affected by the hematocrit level, the glucose concentration was changed within the error range of 5%. This is because the current was kept stable because the red blood cells and white blood cells were completely blocked by the fibers.

As described above, although the embodiments have been described with reference to the limited embodiments and drawings, various modifications and variations are possible from the above description by those skilled in the art. For example, even if the described techniques are performed in an order different from the described method, and/or the described components are combined or combined in a different form from the described method, or replaced or substituted by other components or equivalents Appropriate results can be achieved.

Therefore, other implementations, other embodiments, and equivalents to the claims are also within the scope of the following claims.

Claims (13)

a substrate including a plurality of electrodes; and
a porous matrix layer;
The porous matrix layer includes glucose oxidase-coated polyvinyl alcohol (PVA)/polyacrylic acid (PAA) nanofibers, and the PVA and PAA molar ratio of the nanofibers is 2:3 to 3:2, which is a biosensor for glucose measurement.
delete According to claim 1,
The glucose oxidase-coated nanofiber has a diameter of 400 to 600 nm, a biosensor.
According to claim 1,
In the porous matrix layer, the pore diameter before and after adding the sample is maintained at 2.5 to 5.0 μm, the biosensor.
According to claim 1,
The thickness of the porous matrix layer is 10 to 30 ㎛, biosensor.
According to claim 1,
In the biosensor, the glucose concentration measured in an environment of a hematocrit level of 35% to 60% is constantly measured within an error range of 5%, the biosensor.
7. The method of claim 6,
The glucose concentration, which is constantly measured within an error range of 5% in the biosensor, is 37.1 mg/dL to 544.7 mg/dL, the biosensor.
According to claim 1,
The biosensor, the biosensor, wherein the glucose concentration is measured in 1 to 3 seconds from the time the sample is introduced into the porous matrix.
preparing a substrate including a plurality of electrodes;
Preparing a hydrophilic polymer solution having a molar ratio of 2:3 to 3:2 of polyvinyl alcohol (PVA) and polyacrylic acid (PAA);
Electrospinning the hydrophilic polymer solution on the substrate to prepare a porous matrix made of nanofibers;
heat-treating the porous matrix; and
Coating the porous matrix with glucose oxidase; comprising, a biosensor manufacturing method for glucose measurement.
delete 10. The method of claim 9,
The step of electrospinning the hydrophilic polymer solution on the substrate,
The concentration of the hydrophilic polymer solution is 8 to 20 wt%,
The voltage of the electrospinning is 1 to 80 kV,
The spinning rate is 5 to 30 ml/hr, and
The radiation distance is 10 to 30 cm, the biosensor manufacturing method.
10. The method of claim 9,
The step of coating the porous matrix made of the nanofibers with glucose oxidase,
A method for manufacturing a biosensor, which is performed using a screen trapping method, a bubble impregnation method, a spray method, or a dispensing method.
10. The method of claim 9,
In the porous matrix layer, the pore diameter before and after the sample is added is maintained at 2.5 to 5.0 μm, the biosensor manufacturing method.

Images